In general, the further an object is from the centre of an optical system's optical axis, the worse the imaging system's Strehl ratio is and thus the dimmer the object will appear and the worse will be its signal to noise ratio. Moreover, there is a trade-off between field of view and numerical aperture: the higher the numerical aperture of an optical system is (i.e. the smaller the object it can resolve), the smaller is the field of view over which the system's Strehl ratio is acceptably large (particularly for in vivo systems), and over which vignetting is acceptably small. This trade-off arises from the change of the relative geometry of the imaging optics, the object and the object's image as the position of an object changes in the field of view. Practical optical systems that are optimized for on-axis imaging are not simultaneously optimized for off-axis imaging. The higher the system's numerical aperture, the greater the sensitivity of any design to a change in relative geometry between object, imaging system and image. Aberrations, particularly coma, increase off-axis. The focal position at the drive and fluorescence wavelengths must be the same to within the lateral and axial resolutions of the instrument.
For example, FIG. 1 is a schematic view of a many-element confocal collector lens of the background art, with a 0.3 NA, 0.8 Strehl ratio, axial chromatic shift <2 mm and lateral chromatic shift <150 nm over the whole field of view (FOV). In order to manufacture this lens with a Strehl ratio of 0.5 or greater throughout the whole field of view, it is necessary to consider the limitations of existing manufacturing techniques, and the aggregation of manufacturing imperfections, etc., and to take into account the expected radius of curvature of the imaging surface in the tissue (in this example, >3 mm). Hence, a design or theoretical Strehl ratio of 0.95 is required, at both the drive wavelength of 488 nm and fluorescence emission peak wavelength of 532 nm, together with a lateral focal chromatic shift of less than 150 nm and an axial focal chromatic shift of less than 2 μM between the two wavelengths across the whole field of view, all with the radius of curvature of the imaging surface in the tissue being greater than 3 mm. This is achieved by using the multi-element design of lens shown in FIG. 1.
The concatenation of ten optical elements in turn makes it costly to achieve the required stringent optical quality and mechanical tolerances. Assuming that the aberration from imperfections of each surface add incoherently, the aberration of the fourteen air to glass interfaces is roughly √{square root over (14)}≈3.7 times the aberration added by each surface. The required ISO 10110 specifications for each surface within the final assembly is 3/0.5(0.5/−) RMSi <0.05, λ=633 nm and 4/5′.
A further problem with these systems is the low numerical aperture of the output field of the fibre. The numerical aperture (NA) of the field output from a single mode 450 nm fibre is approximately 0.1. For many clinical in vivo imaging applications, numerical apertures of 0.2 or greater are needed. Therefore, optical magnification is needed if a single mode optical fibre's bound eigenfield is to be used as the confocal pinhole to boost this numerical aperture by a factor of two or more. This means that the scanning amplitude of the fibre must be two or more times that of the field of view in the tissue. A scanning system with a magnification of 2× must achieve a scan amplitude of the optical fibre twice that of the wished-for field of view.
A system with a simpler optical arrangement is disclosed in US Patent Application Publication No. 2011/0211104 and further explored in “High-resolution resonant and nonresonant fiber-scanning confocal microscope”, J. Biomedical Optics 16(2), 026007 (February 2011). US 2011/0211104 discloses an optical probe for a confocal scanning endoscope. The probe comprises an optical guide, a first lens mounted on a distal end portion of the optical guide for focusing light from the optical guide, an actuator for displacing the distal end portion and the first lens to enable optical scanning, and a second lens inside the probe to receive radiation from the first lens. The second lens, which comprises a negative lens, deflects radiation from the first lens in a direction corresponding to a direction of displacement of the first lens by the actuator. The invention is said to be particularly useful for increasing the field of view (FOV) of cheap, disposable optical probes. Thus, the first lens is mechanically coupled to the optical guide, which avoids the trade-off between field of view and numerical aperture, permitting high values of both parameters, and also eliminates lateral chromatic shift as the relative geometry of the lens, imaged tissue and coverslip stays the same as the on-axis geometry throughout scanning.
However, the system of US 2011/0211104 includes several electrical channels running along a scanning steel tube that couples the first lens to the optical guide, employs large, heavy lenses and does not address the problem of chromatic aberration (other than to calculate the pulse spread for many-photon imaging).